1. Field of the Invention
The present invention relates to devices and methods for measuring one or more parameters of a biological sample. More specifically, this invention relates to devices and methods for non-invasively detecting the presence or concentration of one or more analytes in vivo by Raman scattering.
2. Discussion of the Art
When light impinges on a sample, most of the scattered photons are elastically (or Rayleigh) scattered, meaning that they have the same frequency (or wavelength) as the incident radiation. A small fraction of the scattered light (less than one in a thousand incident photons) is inelastically (or Raman) scattered at frequencies that differ from the incident frequency by a value determined by the molecular vibrations of the sample. Raman scattering occurs at frequencies corresponding to the incident frequency plus or minus a molecular vibrational frequency as shown in equation (1): EQU .nu..sub.Raman =.nu..sub.o .+-..nu..sub.vib (1)
where .nu..sub.Raman represents the Raman scattered frequency, .nu..sub.o represents the frequency of incident light (laser), and .nu..sub.vib represents a vibrational frequency of the molecule under study. A Raman spectrum is thus a plot of the intensity of scattered light as a function of frequency (or wavelength). By convention, Raman spectra are reported using wave number values (reciprocal centimeters) so that the abscissa is linear in energy.
Raman spectra have been used to characterize the structure and function of biological molecules, and in some cases, to identify and quantify the composition of complex, multicomponent samples. There are a number of recognized advantages to Raman spectroscopy as an analytical technique. It provides vibrational spectra that are rich in highly reproducible, detailed features, thereby providing the possibility of highly selective determinations. In comparing Raman and infrared (IR) spectroscopies, the Raman approach is advantageous for several reasons: (1) aqueous solutions present no special problems, (2) the low frequency region is easily obtained, and (3) optical components, such as windows, sample containers, and optical fibers can be made out of relatively inexpensive and readily available materials. Compared to absorption techniques, the Raman approach also permits considerable flexibility in the physical state of the sample. Because the selection rules for Raman scattering are different from those of IR absorbance, Raman scattering is complementary to IR measurements. In other words, vibrational modes that produce intense Raman bands may be invisible in the IR spectra.
Several inventors have recognized the potential for using Raman scattering as a non-invasive (NI) sensor. As defined herein, a "non-invasive" sensor is one that can be used without removing a sample from, or without inserting any instrumentation into, the tissues. U.S. Pat. No. 5,553,616 discloses the use of Raman scattering with excitation in the near infrared (780 nm) and an artificial neural network for measuring blood glucose. WO 92/10131 discusses the application of stimulated Raman spectroscopy for detecting the presence of glucose. Co-pending application Ser. No. 08/982,839, filed Dec. 2, 1997, incorporated herein by reference, describes a noninvasive glucose sensor that combines Raman measurements with complementary non-invasive techniques in order to enhance the sensitivity and selectivity of the measurement.
A major challenge for all of the NI Raman techniques to date has been to collect spectral information with sufficiently high signal-to-noise ratios to discriminate weak analyte signals from the underlying background. NI Raman measurements are hindered by a number of factors, some of which are listed below.
(a) Low Quantum Efficiency of Raman Scattering
The quantum efficiency of "normal" Raman scattering is low, thereby necessitating long integration times and high power densities to achieve acceptable signal-to-noise (S/N) ratios. Resonance and surface enhancement offer possibilities for extending the sensitivity and selectivity of Raman measurements and the intensity of some bands can be enhanced by as much as 10.sup.6 over the normal Raman spectrum. However, these enhancement mechanisms are not generally applicable to all analytes or to all samples, and relating band intensities to analyte concentrations under such circumstances requires careful calibration procedures.
(b) Fluorescence Background
In the visible region of the electromagnetic spectrum, indigenous tissue chromophores produce a broad and sloping fluorescence background that is difficult to subtract. Baselines encountered in Raman spectroscopy usually vary from one spectrum to the next even when samples are nominally the same. For in vivo measurements, both the intensity and the direction of the slope can be expected to change and Raman peaks are usually superimposed on slightly curved baselines. Fluorescence signals can be reduced by shifting to near-IR excitation, at the expense of reduced Raman intensity. Time resolved measurements, which discriminate against the fluorescence based on the differing temporal behaviors of the fluorescence and Raman signals, have also been used with some success.
(c) Spectral and Physiological Variables
In the ideal case, a NI Raman sensor would be highly sensitive for the parameter of interest (e.g., analyte concentration) while remaining insensitive to interfering analytes or physiological parameters. In practice, all of the NI Raman techniques described in the prior art are sensitive to one or more interfering "physiological" or "spectral" variables.
As used herein, the expression "physiological variables" describes physiological parameters, such as temperature or pulsatile blood flow, that can adversely affect the sensitivity or selectivity of a NI Raman measurement. Examples of several important physiological variables are listed in Table 1 below. As used herein, the expression "spectral variables" describes spectral features that arise either from poorly resolved analyte bands or from other interfering components in the sample. Several significant sources of spectral interference in biological samples, such as hemoglobin, albumin, cholesterol, urea, and fat, are listed in Table 2 below. Other tissue constituents that are present at lower concentrations or have lower scattering cross-sections may also contribute to an overall background signal that is difficult to subtract.
Referring again to Table 1 and Table 2, it is important to note that each of the physiological and spectral variables may fluctuate over time and each variable may oscillate at a different frequency. Methods that do not account for these oscillating variables will provide inaccurate results.
(d) Self-absorption
When a Raman band of the analyte falls within an absorption band of one of the sample components, much of the Raman scattered light may be absorbed by the sample, resulting in an underestimate of the analyte concentration. The degree of self-absorption is highly dependent upon the geometry of the measurement (i.e., orientation of the optics) and the sample composition.
(e) Instrument Size and Complexity
The vast majority of Raman scattering measurements in the prior art have been (and continue to be) performed in experimental research laboratories with the aim of elucidating the molecular structure, bonding, or reactivity of constituent molecules in a sample. Accordingly, the instruments used for most Raman measurements (i.e., dispersive Raman spectrometers) are designed to maximize experimental accuracy and flexibility. Most Raman spectrometers employ dispersive devices, such as Czerny-Turner spectrometers, to spread the scattered light spatially into a spectrum of different wavelength components. The Czerny-Turner spectrometer can be operated in two different modes, as described below.
First, the spectrometer can be used as a monochromator (i.e., selecting only one wavelength for analysis). For single wavelength detection, the spectral bandpass is determined by the entrance and exit slits of the monochromator and the resolution of the diffraction grating employed therein. The bandpass of the spectrometer can be adjusted to the spectral resolution required by the measurement (typically 0.5-10 cm-1). As used herein, the bandpass describes the full width of the band at half its maximal intensity (FWHM). The spectrometer diffraction grating is rotated (scanned), and a discrete detector, such as a photomultiplier tube, measures the intensity of the scattered light transmitted by the spectrometer.
Alternatively, the Czerny-Turner spectrometer can be used as a spectrograph, whereby a range of wavelengths is selected. In this case, an array detector, such as a photodiode array or an area detector, such as a charge coupled device (CCD), is placed at the back focal plane of the spectrometer (in place of the exit slit). Spectral resolution is controlled by the entrance slit of the spectrometer and the groove density of the diffraction grating. The scattered intensity is recorded over a preselected wavelength range, thereby providing a multi-channel advantage over the scanning method described above. However, the grating may require movement and recalibration, however, if the wavelength range of the measurement is changed.
Although Czerny-Turner spectrometers provide highly accurate and adjustable selection of the wavelength and spectral bandpass, the transmissivity of these devices is low. Thus, the time required to obtain a complete Raman spectrum is unacceptably long for use as a NI sensor. Although the instrument can provide measurements of discrete spectral features in the scanning mode, only one wavelength may be sampled at any given time. Further, valuable time is lost while the grating is moved to interrogate different portions of the Raman spectrum. The mechanical motion required by this system can also cause repositioning errors, thereby increasing uncertainty of accuracy of wavelength. In general, devices employing Czerny-Turner spectrometers are too large, complicated, and expensive for field use by inexperienced operators.
Recent progress in the area of electronically tunable filters offers some hope for reducing the size of Raman spectral filtering devices. The acousto-optic tunable filters (AOTFs) can provide broad tunability and multiplex operation. However, the spectral bandpass of an AOTF is too large for most Raman measurements. A liquid crystal tunable filter (LCTF) can provide a narrower spectral bandpass, but such a filter is limited to the detection of one wavelength at a time. The maximum transmissivity of a LCTF (approximately 16%) also limits the sensitivity of this device.
Fixed filters are a cost-effective alternative to the dispersive devices and electronically tunable filters described above. Although fixed filters lack the flexibility of the variable wavelength selective devices, they are robust, compact, and much less expensive than dispersive instruments or electronically tunable filters. Their high transmissivities (&gt;80%) reduce signal acquisition times, and their spectral properties (e.g., .lambda..sub.max, .DELTA..nu..sub.1/2, rejection ratio) are selected at the factory, thereby eliminating the need for electronic controllers and moving parts.
Fixed filters are designed to attenuate all but the desired wavelengths of radiation. For example, they can be used to pass a band of wavelengths (bandpass filters) or to block wavelengths longer or shorter than some desired value (cutoff filters). Bandpass filters are characterized by a plot of their spectral transmittance vs. wavelength, as shown in FIG. 1(a), which exhibits a characteristic wavelength of maximum transmission (.lambda..sub.max) and a bandpass (.DELTA..nu..sub.1/2) measured as the full width at half maximum. Most bandpass filters are made with alternating layers of high-refractive and low-refractive index dielectrics. Multilayer interference filters can achieve quite narrow FWHM bandpass (less than 10 cm.sup.-1) values with high peak transmittances (greater than 50%) from the ultraviolet to the infrared.
Cutoff filters attenuate radiation at wavelengths shorter than a given cutoff wavelength (short-wavelength cutoff) or radiation at longer wavelengths than a given cutoff wavelength (long-wavelength cutoff). The transmittance of a short-wavelength cutoff filter is illustrated in FIG. 1(b). Dichroic mirrors are wavelength selective beam splitters with transmission characteristics that are similar to cutoff filters. Longpass dichroic mirrors transmit radiation at wavelengths that are longer than a given cutoff wavelength and reflect radiation at shorter wavelengths.
Holographic filters have the sharp cut-off characteristics of the multilayer dielectric filters described above and high transmissivities (80-90%). The transmission curves of holographic filters are featureless, as opposed to the transmission curves of dielectrics, which have numerous features due to multiple layer interferences.
EP 0781990 to Ozaki et al., hereinafter Ozaki '990, describes a Raman filtering device for urinalysis comprising multiple bandpass filters or a combination of cutoff filters for selecting particular wavelengths of Raman scattered light. For measurements at multiple wavelengths, Ozaki '990 provides a filter wheel, which can be rotated to provide several different filters in succession.
Ozaki '990 is inappropriate for in vivo measurements where spectral and physiological variables fluctuate over time. In particular, the device of Ozaki '990 is incapable of collecting Raman signals at two wavelengths simultaneously. Changing filters by spinning the filter wheel also causes slight changes in alignment, which can alter collection efficiency, thereby providing erroneous measurements of in vivo analyte concentrations.
Despite the wide variety of instrumental configurations used to date, NI Raman sensors are not yet commercially available. Accordingly, there is a need for a compact, rugged, inexpensive, sensitive, and selective Raman sensor that measures in vivo analyte concentrations in the presence of fluctuating physiological and spectral variables.